In the last few years advances in nuclear magnetic resonance (NMR) techniques have made it possible to form two and three dimensional spin density images of solids and liquids. A number of novel and sophisticated variants have also been introduced to the rapidly expanding field of imaging. An important aspect of all these developments is the ability to form images of biological tissue in vivo. The NMR method is non-invasive and has a much lower radiation hazard than the more usual X-ray imaging methods.
In addition to producing spin density pictures, these new NMR imaging techniques can all be adapted to measure spatial variations of the spin-lattice relaxation time in a specimen. The cell water in cancerous tissue, for example, is known to have longer spin-lattice relaxation time than that in normal tissue. Thus NMR imaging, though in its infancy, holds promise as a diagnostic tool for the early detection of tumors.
The NMR imaging techniques to be described all rely on the preparation and/or observation of the nuclear spin system in the presence of one or more magnetic field gradients. The field gradients serve to spatially differentiate regions of the specimen by changing the Larmor resonance frequency of the spins from one region of the specimen to another.
Individual protons or hydrogen nuclei are found in most organic and biological material and have a natural isotopic abundance of 99.9844 percent. The other 0.0156 percent of nuclear sites is taken up with the other naturally occurring heavy hydrogen isotope, deuterium.
Each nucleus has associated with it a small nuclear magnetic moment and a quantity of angular momentum called spin. Regarded classically, the combined effect of magnetic moment and spin causes a proton to precess about the direction of an applied static magnetic field much as a spinning top precesses about the gravitational field direction if perturbed from the upright position. For protons, the precessional frequency is independent of the angle of inclination of the magnetic moment with respect to the static magnetic field and is called the Larmor angular frequency .omega..sub.o. However, it does depend directly on the magnitude of the static magnetic field B.sub.o through the relationship .omega..sub.o =.gamma.B.sub.o, where the constant .gamma. is called the magnetogyric ratio. This relationship is the key to much of what follows. If B.sub.o is varied then .omega..sub.o will vary. If a linear magnetic field gradient is superimposed on an otherwise spatially uniform B.sub.o, then the protons in a specimen placed in these fields would experience a magnetic field higher than B.sub.o in some places and lower in others. An account of the development of NMR spin imaging is given by Mansfield, Contemp. Phys. Vol. 17, No. 6, pp. 553-576 (1976). The basic principles of NMR necessary to understand imaging are discussed and main methods of imaging are described and illustrated with examples of images of proton spin distributions in a number of biological specimens.
NMR imaging of humans for medical diagnostic purposes presents the magnet designer with formidable problems of an unusual nature. Hoult et al., Rev. Sci. Instrum. 52(9), pp. 3142-51 September (1981) state that a magnet is required which produces a field of at least 0.1 T with a homogeneity of prefereably 1 ppm over the region of interest of the patient, say the head or torso. In addition, Hoult et al. state that linear field gradients of up to 10.sup.-2 Tm.sup.-1 in any direction may be required. A short term field stability of better than 0.1 ppm may be mandatory over a period of a second in order to avoid phase noise on the NMR signal, while the long term stability may need to be about 1 ppm. Further, all this must be accomplished in a hospital environment where it is likely that serious perturbations will be caused by large amounts of steel (reinforcement, water pipes, etc.), in the building structure, (elevators, beds, nearby trucks, etc.). Weight unfortunately precludes the use of an iron magnet with its convenient flux return path, and current designs are therefore air-cored electromagnets of either resistive or superconducting design. A spherical shaped electromagnet for NMR imaging is described by Hoult et al., supra. A superconductive NMR magnet for in vivo imaging is described by Goldsmith et al., Physiol. Chem. & Phys., 9, pp. 105-107 (1977). In addition, Hanley discussed superconducting and resistive magnets in NMR scanning in a paper presented at the 1981 International Symposium on Nuclear Magnetic Resonance Imaging held at the Bowman Gray School of Medicine, Wake-Forest University, Winston-Salem, N.C. A superconducting magnet can attain much higher fields than a simple electromagnet but its cost will be much higher.
To date, for various reasons no one has made a permanent magnet NMR apparatus for imaging biological tissue. A permanent magnet system would be superior to prior art superconducting and resistive electromagnet designs in the following ways:
(a) There is no need for a means of generating the large amounts of power required to maintain the field as in the resistive magnet systems. PA1 (b) There is no need to provide cooling means to either remove generated heat as in the resistive magnets or to maintain cryogenic temperatures as in the superconducting magnets. PA1 (c) The field of the permanent magnet is not subject to power supply drift like that of resistive magnets, or superconducting magnets not operated in the persistent current mode. PA1 (d) The field of the permanent magnet is not subject to gradual decay like that of superconducting magnets operating in the persistent mode. PA1 (e) The material used can be a readily available ferrite magnet material that is transparent to electromagnetic waves of the frequencies of interest (5 MHz to 15 MHz). PA1 (f) The external field strength falls off rapidly with distance away from the magnet, leading to significantly reduced interference with the bias field from ferromagnetic objects in the vicinity of the apparatus.